Medical radionuclide imaging, commonly referred to as nuclear medicine, is a unique specialty wherein ionizing radiation is used to acquire images which show the function and anatomy of organs, bones or tissues of the body. The technique of acquiring nuclear medicine images entails first introducing biologically appropriate radiopharmaceuticals into the body—typically by injection, inhalation, or ingestion. These radiopharmaceuticals are attracted to specific organs, bones or tissues of interest (These exemplary organs, bones, or tissues are also more generally referred to herein using the term “objects”). Upon arriving at their specified area of interest, the radiopharmaceuticals produce gamma photon emissions which emanate from the body and are then captured by a scintillation crystal. The interaction of the gamma photons with the scintillation crystal produces flashes of light which are referred to as “events.” Events are detected by an array of photodetectors (such as photomultiplier tubes) and their spatial locations or positions are then calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body. Known applications of nuclear medicine include: analysis of kidney function, imaging blood-flow and heart function, scanning lungs for respiratory performance, identification of gallbladder blockage, bone evaluation, determining the presence and/or spread of cancer, identification of bowel bleeding, evaluating brain activity, locating the presence of infection, and measuring thyroid function and activity. Hence, accurate detection is vital in such medical applications.
One particular nuclear medicine imaging technique is known as positron emission tomography, or PET. Positron emission tomography is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. The measurement of tissue concentration using a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from a positron annihilation. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by an annihilation event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Annihilation events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors; i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line-of-response (LOR) along which the annihilation event has occurred.
After being sorted into parallel projections, the LOR defined by the coincidence events are used to reconstruct a three-dimensional distribution of the positron-emitting radionuclide within the patient. In two-dimensional PET, each 2D transverse section or “slice” of the radionuclide distribution is reconstructed independently of adjacent sections. In fully three-dimensional PET, the data are sorted into sets of LOR, where each set is parallel to a particular detector angle, and therefore represents a two dimensional parallel projection p(s, φ) of the three dimensional radionuclide distribution within the patient—where “s” corresponds to the distance of the LOR from the center of the detector and “φ” corresponds to the angle of the detector plane with respect to the x axis in (x, y) coordinate space (in other words, φ corresponds to a particular LOR direction).
Coincidence events are integrated or collected for each LOR and stored in a sinogram. In this format, a single fixed point in f(x, y) traces a sinusoid in the sinogram. In each sinogram, there is one row containing the LOR for a particular azimuthal angle φ; each such row corresponds to a one-dimensional parallel projection of the tracer distribution at a different coordinate along the scanner axis. This is shown conceptually in FIG. 1.
An event is registered if both crystals detect an annihilation photon within a coincidence time window τ (e.g., on the order of 4-5 nsec), depending on the timing properties of the scintillator and the field of view (FOV). The FOV is defined as the volume between the detectors; and a pair of detectors is sensitive only to coincidence events occurring in the FOV. Therefore, the need for physical collimation is eliminated and sensitivity is significantly increased. Accurate corrections (for example, attenuation correction) can be made for the self-absorption of photons within the patient so that accurate measurements of tracer concentration can be made.
The number of time coincidences detected per second within a FOV of a detector is the count rate of the detector. The count rate at each of two oppositely disposed detectors, A and B, can be referred to as singles counts or SA and SB, respectively. The time required for a gamma photon to travel from its point of origin to a point of detection is referred to as the time-of-flight (TOF) of the gamma photon. TOF is dependent upon the speed of light c and the distance traveled. A time coincidence or coincidence event is identified if the time difference between the arrivals of signals in a pair of oppositely disposed detectors is within the coincidence time window τ. In conventional PET, the coincidence detection time window τ is wide enough so that an annihilation event occurring anywhere within the object will produce annihilation gamma photons reaching their respective detectors within the coincidence window. Coincidence time windows of 4.5-12 nsec are common for conventional PET, and are largely determined by the time resolution capabilities of the detectors and electronics.
As illustrated in FIG. 2, if an annihilation event occurs at the midpoint of a LOR, the TOF of the gamma photon detected in detector A (TA) is equal to the TOF of the gamma photon detected in detector B (TB). If an annihilation event occurs at a distance Δx from the midpoint of the LOR, the difference between TA and TB is Δt=2Δx/c, where c is the speed of light. If d is the distance between detectors, the TOF difference Δt could take any value from −d/c to +d/c, depending on the location of the annihilation event.
Time-of-flight (TOF) positron emission tomography (PET) (“TOF-PET”) is based on the measurement of the difference Δt between the detection times of the two gamma photons arising from the positron annihilation event. This measurement allows the annihilation event to be localized along the LOR with a resolution of about 75-120 mm FWHM, assuming a time resolution of 500-800 ps (picoseconds). Though less accurate than the spatial resolution of the scanner, this approximate localization is effective in reducing the random coincidence rate and in improving both the stability of the reconstruction and the signal-to-noise ratio (SNR), especially when imaging large objects. Thus, in TOF-PET, the “TOF” coordinate, Δt, is stored together with s and φ.
One task of a good PET detector is to measure the time stamp of the gamma interaction in the detector (usually within the scintillator) as accurately as possible. The required accuracy of a few hundred ps (picoseconds) is on a much shorter time scale than the width or duration of the photon emission and also, normally, the width of the pulse from the photosensor for a single photon. The figure of merit is the coincidence time resolution, which is defined as the full width half maximum (FWHM) of the measured trigger time difference between two detectors, when detecting the signal from a fixed point source positron emitter between them.
Time resolution is a measure of the uncertainty of the measured time difference Δt between the two detections. The time resolution is used in the reconstruction algorithm as a kernel for a localization probability function. The events are located along the LOR identified by the two detectors; its most probable position is set to the position corresponding to the measured TOF difference t.
There are various types of photosensors that are used or have been studied for use in PET scanners: most commercial systems today are based on photomultiplier tubes (PMTs), which detect the light emitted by a scintillator following a gamma event and convert detected light photons into electrical signals. Some recent systems and scanner prototypes have used avalanche photodiodes (APDs) to detect the gamma event. Silicon photomultipliers (SiPMs) seem to be a promising sensor type for future detector generations, because they combine some advantages of PMTs (high signal gain and high speed) with those of APDs (small form factor and compatibility with magnetic fields).
SiPMs offer some of the best time resolutions among existing photosensor devices, and in particular are very suitable for detecting the scintillation light from Lutetium Oxyorthosilicate (LSO) and other typical PET scintillators. Coincidence time resolution values in the range of 150 to 500 ps have been obtained for different sensor types and different coupling configurations, demonstrating the feasibility of this technology for TOF-PET. The best timing resolution is usually achieved by coupling each single LSO crystal with typical dimensions of 3×3×20 mm3 to a single sensor pixel, which is matched to the 3×3 mm2 light extraction face.
Analog SiPMs consist of an array of Geiger-mode APDs (or microcells) connected in parallel to form a two-terminal device. Then, although the state of the individual microcell can be described as digital (ON or OFF), the overall output becomes an analog signal, which is roughly proportional to the amount of incident light.
One very recent development in this field is the so-called digital SiPM (dSiPM) technology by Philips (see, e.g., WO2009/019660 A2). This sensor design achieves a high signal-to-noise ratio and an excellent time resolution by integrating CMOS electronics on the same wafer as the SiPM sensor and digitizing the signal of each Geiger-mode APD cell. In contrast to the analog sensors discussed above, the dSiPM can directly measure the trigger time for the first, second, third, . . . avalanche, thus giving measured arrival times for the first few measured photons. Setting the trigger on the first photon has been shown to give the best timing results for any SiPM so far (T. Frach et al., “Digital Silicon Photomultiplier—Principle of Operation and Intrinsic Detector Performance,” IEEE Nuclear Science Symposium, Talk 28-5, 2009). It has been shown theoretically that the first few photons from the emission statistics carry the most precise information about the time of the original event interaction (Y. Shao, “A new timing model for calculating the intrinsic timing resolution of a scintillator detector,” Phys. Med. Biol. 52 (2007) 1103-1117).
Increasing the accuracy of the trigger time for detecting the onset of a gamma interaction event is especially important for time-of-flight PET (TOF-PET), where the variance in the reconstructed image is reduced by using the additional spatial information from the trigger time differences of the two opposite detectors that measure the same positron decay (W. Moses, IEEE Trans. Nucl. Sci. vol NS-50, p. 1325-1330, 2003).
In time-of-flight PET, the noise equivalent count rate (a measure of the detected counts corrected for the noise contribution of scatter and random coincidences; in other words, a measure of the effective sensitivity of the PET scanner) is inversely proportional to the time resolution of the PET scanner. Thus, improving the time resolution is the key for better performing TOF-PET scanners. Table 1 illustrates the time resolution, spatial uncertainty, and estimated TOF noise equivalent count rate (NEC) gain for a 40 cm diameter uniform cylinder (M. Conti, “State of the art and challenges of time-of-flight PET,” Physica Medica 25: 1, p. 5, 2008).
TABLE 1Effects of Improved Time ResolutionTime resolution (ns)Δx (cm)TOF NEC gain0.11.526.70.34.58.90.57.55.51.218.02.2